Method of fabricating an implantable medical device by controlling crystalline structure

ABSTRACT

A method of fabricating an implantable medical device that includes deforming and heating setting a polymer construct, for use in fabricating the device, in a temperature range in which the crystal nucleation rate is greater than the crystal growth rate is disclosed.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to a method of fabricating an implantable medicaldevice by controlling crystalline structure.

2. Description of the State of the Art

This invention relates generally to implantable medical devices having arange of mechanical and therapeutic requirements during use. Inparticular, the invention relates to radially expandable endoprosthesesthat are adapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen subjected to angioplasty or valvuloplasty.

The stent must be able to satisfy several mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel lumen. This requires a sufficient degree ofstrength and rigidity or stiffness. In addition to having adequateradial strength, the stent should be longitudinally flexible to allow itto be maneuvered through a tortuous vascular path and to enable it toconform to a deployment site that may not be linear or may be subject toflexure. The material from which the stent is constructed must allow thestent to undergo expansion which typically requires substantialdeformation of portions of the stent. Once expanded, the stent mustmaintain its size and shape throughout its service life despite thevarious forces that may come to bear thereon, including the cyclicloading induced by the beating heart. Therefore, a stent must be capableof exhibiting relatively high toughness which corresponds to highstrength and rigidity, as well as flexibility.

A stent is typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts. The stent canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. A pattern can be formed in a tube, for example, bylaser cutting. The scaffolding is designed to allow the stent to beradially expandable. The pattern is generally designed to maintain thelongitudinal flexibility and radial rigidity required of the stent.Longitudinal flexibility facilitates delivery of the stent and radialrigidity is needed to hold open a bodily lumen.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesa bioactive agent. Polymeric scaffolding may also serve as a carrier ofa bioactive agent.

In many treatment applications, the presence of a stent in a body may benecessary for a limited period of time until its intended function of,for example, maintaining vascular patency and/or drug delivery isaccomplished. Therefore, stents fabricated from biodegradable,bioabsorbable, and/or bioerodable materials such as bioabsorbablepolymers should be configured to completely erode only after theclinical need for them has ended.

A polymeric implantable medical device should be mechanically stablethroughout the range of stress experienced during use of an implantablemedical device. Unfortunately, many polymers used for stent scaffoldingsand coatings are relatively brittle under physiological conditions,e.g., at body temperature. Many polymers remain relatively brittle, andhence susceptible to mechanical instability such as fracturing while inthe body.

What is needed is an implantable medical device capable of satisfyingmechanical requirements during desired treatment time.

SUMMARY

The invention provides a method for fabricating an implantable medicaldevice, the method comprising: deforming a polymer construct whilemaintaining the polymer construct at a temperature range of from aboutTg to about 0.6(Tm−Tg)+Tg; maintaining the deformed polymer construct inthe deformed state at a temperature range of from about Tg to about0.9(Tm−Tg)+Tg for a sufficient period of time to heat set the polymerconstruct; and fabricating an implantable medical device from the heatset polymer construct.

Further, the invention provides a method for fabricating an implantablemedical device comprising: deforming a polymer construct at a firsttemperature greater than a Tg of the polymer at which the crystalnucleation rate of the polymer construct is greater than the crystalgrowth rate; maintaining the deformed polymer construct in the deformedstate at a second temperature greater than the Tg of the polymer atwhich the crystal nucleation rate of the polymer construct is greaterthan the crystal growth rate for a sufficient period of time to heat setthe polymer construct; and fabricating an implantable medical devicefrom the set polymer.

Finally, the invention provides an implantable medical devicecomprising: a plurality of crystalline domains having crystals dispersedwithin an amorphous domain, the majority of the crystals being less thanabout 2 microns.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 depicts a three-dimensional view of a stent which is made up ofstruts.

FIG. 2 depicts a graph illustrating the nucleation rate and growth rateof crystals in a polymer versus temperature of a polymer.

FIG. 3A depicts a portion of a strut shown in FIG. 1 of a polymer tubedeformed and heat set at temperatures closer to Tm than Tg according tothe invention.

FIG. 3B depicts a schematic representation of the microstructure of apolymer tube deformed and heat set at temperatures closer to Tm than Tg,having fewer crystals of a larger size compared to the polymeric tube inFIG. 4B, which is deformed and heat set at a temperature closer to Tgthan Tm.

FIG. 4A depicts a portion of a strut shown in FIG. 1 of a polymer tubedeformed and heat set at temperatures closer to Tg than Tm according tothe invention.

FIG. 4B depicts a schematic representation of the microstructure of apolymer tube deformed and heat set at temperatures closer to Tg than Tmaccording to the invention, having fewer crystals of smaller sizecompared to the polymer tube in FIG. 3B, which is deformed and heat setat a temperature closer to Tm than Tg.

DETAILED DESCRIPTION

The embodiments of the present invention relate to implantable medicaldevices and methods to control the relationship between degree ofnucleation and associated crystal growth to improve mechanicalproperties such as strength and flexibility, as well as an extendedshelf life.

For the purposes of the present invention, the following terms anddefinitions apply:

As used herein, “polymer construct” refers to any useful article ofmanufacture made of a polymer such as a semi-crystalline polymer, orblend of polymers, any useful article of manufacture made of anymaterial that is coated with a polymer or blend of polymers. Someexamples of polymer constructs include, but are not limited to, a tube,a sheet, a fiber, etc.

“Glass transition temperature,” T_(g), is the temperature at which thepolymer's amorphous domains transform from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the T_(g) corresponds to the temperature where segmental motionstarts in the polymer chains. When an amorphous or semicrystallinepolymer is exposed to an increasing temperature, both the polymer'scoefficient of expansion and the heat capacity increase as thetemperature is raised, indicating increased molecular motion. As thetemperature is raised, the actual molecular volume in the sample remainsconstant, and so a higher coefficient of expansion points to an increasein free volume associated with the system and therefore increasedfreedom for the molecules to move. The increasing heat capacitycorresponds to an increase in heat dissipation through movement. Tg of agiven polymer can be dependent on the heating rate and can be influencedby the thermal history of the polymer. Furthermore, the chemicalstructure of the polymer heavily influences the glass transition byaffecting mobility.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness are energy perunit volume of material. See, e.g., L. H. Van Vlack, “Elements ofMaterials Science and Engineering,” pp. 270-271, Addison-Wesley(Reading, Pa., 1989).

A brittle material is a relatively stiff or rigid material that exhibitslittle or no plastic deformation. As stress is applied to a brittlematerial, it tends to fracture at a stress approximately equal to itsultimate strength, undergoing little or no plastic deformation in theprocess. A polymer below its Tg tends to be brittle. In contrast, aductile material under an applied stress exhibits both elastic andplastic deformation prior to fracture. Above its Tg, a polymer isductile.

A fracture may be categorized as either ductile or brittle. A relativelylow amount of energy is required to fracture brittle materials.Conversely, ductile materials can absorb a relatively high amount ofenergy prior to fracture. Therefore, ductile materials tend to exhibithigher toughness than brittle materials. Toughness is a desirablecharacteristic in implantable medical devices.

The contrast between brittle fracture and ductile fracture is importantin relation to an implantable medical device because it is useful incharacterizing the device. The mechanical limits of an implantablemedical device during use can be more accurately characterized by theamount of energy absorbed by the device rather than the strength of thedevice material. For example, two devices may be made of differentmaterials having the same or similar ultimate strength. However, underthe same conditions, a material with a lower toughness will fail beforea material with a higher toughness.

As used herein, an “implantable medical device” refers to any type ofappliance that is totally or partly introduced into a patient's body,and which is intended to remain there after the procedure. Examples ofimplantable medical devices include, without limitation, self-expandablestents, balloon-expandable stents, stent-grafts, implantable cardiacpacemakers and defibrillators; leads and electrodes for the preceding;implantable organ stimulators such as nerve, bladder, sphincter anddiaphragm stimulators, cochlear implants, artificial bone; prostheses,vascular grafts, grafts, artificial heart valves and cerebrospinal fluidshunts. Of course, an implantable medical device specifically designedand intended solely for the localized delivery of a therapeutic agent iswithin the scope of this invention. The implantable medical device maybe constructed of any biocompatible material.

FIG. 1 depicts a three-dimensional view of a stent 100. Stent 100includes struts 110, which can take on a variety of patterns. Thestructural pattern of the device can be of virtually any design. Theembodiments disclosed herein are not limited to stents or to the stentpattern illustrated in FIG. 1, but instead can be applied to other stentpatterns and other devices. As shown in FIG. 1, the geometry or shapesof stents vary throughout its structure.

As indicated above, an implantable medical device, such as a stent,should be capable of exhibiting relatively high strength and rigidity,as well as flexibility since devices have varied mechanical requirementsduring use, both before and during treatment.

An implantable medical device may be configured to degrade afterimplantation by fabricating the device either partially or completelyfrom biodegradable polymers. Polymers can be biostable, bioabsorbable,biodegradable, or bioerodable. Biostable refers to polymers that are notbiodegradable. The terms biodegradable, bioabsorbable, and bioerodable,as well as degraded, eroded, and absorbed, are used interchangeably andrefer to polymers that are capable of being completely eroded orabsorbed when exposed to bodily fluids such as blood and may begradually absorbed and eliminated by the body.

A biodegradable device may be intended to remain in the body until itsintended function of, for example, maintaining vascular patency and/ordrug delivery is accomplished. For biodegradable polymers used incoating applications, after the process of degradation, erosion,absorption has been completed, no polymer will remain on the stent. Insome embodiments, very negligible traces or residue may be left behind.The duration is typically in the range of six to twelve months.

Representative examples of polymers that may be used to fabricate, coat,or modify an implantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(L-lactide-co-ε-caprolactone), poly(trimethylenecarbonate), polyester amide, poly(glycolic acid-co-trimethylenecarbonate), co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes,biomolecules (such as fibrin, fibrinogen, cellulose, starch, collagenand hyaluronic acid).

Many polymers used for stent scaffoldings and coatings are relativelybrittle and susceptible to mechanical instability at biologicalconditions. This is particularly true for polymers with a Tg above abody temperature, such as poly L-lactide (“PLLA”), where the polymer inthe stent never reaches its Tg. As a result, PLLA has relatively lowfracture toughness and a relatively low degradation rate at conditionswithin the human body. It is important for a device to have a highfracture toughness throughout the range of stress experienced during useof an implantable medical device.

Other potential problems with polymeric stents are creep, stressrelaxation, and physical aging. Creep refers to the gradual deformationthat occurs in a component subjected to an applied load. Creep occurseven when the applied load is constant. It is believed that the delayedresponse of polymer chains to stress during deformations causes of creepbehavior. Deformation stops when the initially folded chains reach a newequilibrium configuration (i.e. slightly stretched). This deformation isrecoverable after the load is removed, but recovery takes place slowlywith the chains retracting by folding back to their initial state. Therate at which polymers creep depends not only on the load, but also ontemperature (see temperature dependence). In general, a loaded componentcreeps faster at higher temperatures. Long term creep in a polymericstent reduces the effectiveness of a stent in maintaining a desiredvascular patency. In particular, long term creep allows inward radialforces to permanently deform a stent radially inward.

Stress relaxation is also a consequence of delayed molecular motions asin creep. Contrary to creep, however, which is experienced when the loadis constant, stress relaxation occurs when deformation (or strain) isconstant and is manifested by a reduction in the force (stress) requiredto maintain a constant deformation.

Physical aging, as used herein, refers to the densification in theamorphous regions of a semi-crystalline polymer. Densification is theincrease in density of a material or region of a material. Densificationresults from residual and applied stresses. As used herein, a “residualstress” includes, without limitation, the stress in a bulk polymer thatis in a non-equilibrium thermodynamic state. Physical aging ofsemi-crystalline polymers that have glass transition temperatures (Tg)above their normal storage temperature, which, for the purposes of thisinvention is room temperature, i.e., from about 15° C. to about 35° C.,occurs primarily through the phenomenon known as densification.

Densification occurs when a semi-crystalline polymer is cooled at anon-equilibrium rate from a temperature above its T_(g) to a temperaturebelow its T_(g). Such is in fact normally what will occur in mostindustrial settings in that equilibrium cooling is very slow and wouldbe considered economically impractical. The non-equilibrium cooling rateresults in the polymer chains of the amorphous domains being trapped atnon-optimal separation distances in the glassy state that forms when thetemperature goes below T_(g). The chains then attempt to achieve optimalseparation by coordinated localized chain motion. Although thereordering of polymer chains do not result in chain ordering, that is,the formation of lamellae and crystallites, which would constitutecrystallization, the effect on the bulk properties of the polymer issimilar to that obtained when crystallization occurs, that is, themodulus of the polymer increases and concomitantly the polymer becomesmore brittle. Thus, densification of a polymer initially selected forits toughness and elasticity could cause in-use failure of a constructmade of or coated with the polymer when the polymer ages or densifiesand becomes brittle.

Some polymers, such as semi-crystalline polymers usually contain bothamorphous and crystalline domains at temperatures below their meltingpoint. Amorphous regions are those in which polymer chains are situatedin an essentially random orientation. Crystalline domains are those inwhich polymer chains adopt an ordered orientation with segments ofseparate chains or of the same chain becoming essentially parallel toone another to form structures known as lamellae. Lamellae initiallyform from a point of nucleation, which normally is a speck of impurityin the liquid polymer. The formed lamellae then grow outward from thenucleation point to form larger, essentially spherical crystallinestructures known as crystallites.

If the polymer includes inter- or inner-chain crystalline structures,(lamellae and spherulites), which are not capable of movement unlessthey melt the movement of polymer chains in the amorphous domains isreduced and the ability of the chains to densify is correspondinglylessened. For this reason, increasing the degree of crystallinity of apolymer reduces or even eliminates physical aging.

Increased crystallinity, however, can be an undesirable characteristicin polymer constructs in which high toughness is important, sinceincreased crystallinity can confer increased brittleness in a polymer.Specifically, a polymer construct can be brittle in a temperature rangeof use of an implantable medical device. Thus, it would be desirable tofabricate a polymer construct that is sufficiently crystalline tomitigate densification, i.e., physical aging, while reducing oreliminating the changes in bulk properties that accompany increasedcrystallinity, such as increased brittleness. Various aspects of thepresent invention provide a device and a method of fabricating thedevice having high fracture toughness as well as a high enoughcrystallization to reduce physical aging.

It is well known by those skilled in the art that molecular orientationor alignment of polymer chains in a polymer is a particularly importantphenomenon that strongly influences bulk polymer properties such asfracture toughness. Orientation refers to the degree of alignment ofpolymer chains along a longitudinal or covalent axis of the polymerchains. For example, the strength along a particular direction in apolymer is higher when there is alignment of polymer chains along thedirection. Molecular orientation along the preferred direction in apolymeric material may be induced by applying stress along the preferreddirection. Generally, required strength along a particular direction canbe induced in polymer construct for use in fabricating an implantablemedical device.

As mentioned, radial strength is an important characteristic of stents.Therefore, strength and orientation along a circumferential directioncan be induced by radially expanding a tube for fabricating a stent.

Circumferential orientation may be induced through radial expansion of aformed tube. In this process, stress induces orientation of polymerchains, and crystallization is induced by stress and oriented-inducedmechanisms. By deforming or radially expanding a polymer, orientation ofthe polymer chains are induced.

Since polymer chain alignment is a time and temperature dependentprocess, a highly oriented structure that is thermodynamically stable ata given temperature may not be formed instantaneously. Thus, the polymerconstruct may be formed over a period of time. Therefore, afterdeforming a polymer construct, the polymeric construct may be heat set.“Heat setting” refers to allowing aligned polymer chains in the polymerconstruct to equilibrate towards the induced highly oriented structureat an elevated temperature.

In a polymer construct formed in this manner, it is important that theinduced orientation is maintained such that the stent can retain theincreased strength due to the induced orientation. It is desirable forthe stent to be stable as well as have sufficient toughness at thetemperatures of use, e.g., body temperature.

The invention provides for a method and device to control crystallinestructure and maintain induced strength and molecular orientation in adevice. A polymeric material deformed and heat set at the temperaturerange according to the invention forms a greater number of smaller sizedcrystalline domains as compared to polymeric material deformed and heatset a temperature near Tm of the polymer. The greater number of smallersized crystalline domains acts as net points to keep molecularorientation and material stability of the polymer. A polymeric materialhaving a greater number of smaller sized crystalline domains producesmore tie molecules, or polymer chains that link crystalline domains,compared to a small number of larger-sized crystals. A material having agreater number of smaller sized crystalline domains exhibits higherfracture toughness and less brittleness.

A material having a greater number of smaller sized crystalline domainscan be formed by controlling nucleation rate and crystal growth rate,since both are functions of temperature. Nucleation rate is the rate offormation of crystals. Crystal growth rate is the rate at which existingcrystals grow. The invention provides controlling the temperature whiledeforming and heat setting the material such that the degree ofnucleation is larger than the associated crystal growth rate. In thisway, a polymeric material is formed having of a greater number ofsmaller sized crystalline domains and thus increased fracture toughnessand reduced brittleness.

FIG. 2 is a graph of the nucleation rate A and crystal growth rate B asa function of temperature for a polymer is between Tg and Tm. Polymercrystallization begins with nucleation, the formation of smallcrystalline domains that begin as specks of impurities in the amorphousliquid polymer. As depicted, the rate of nucleation A occurs mostrapidly at temperatures near Tg. In contrast to the rate of nucleationA, crystal growth B occurs most rapidly at temperatures near Tm. At atemperature below Tg, there is no increase in crystal growth B orincrease in crystallinity. As the temperature approaches Tm, crystalgrowth B increases logarithmically, at which point the crystals melt andthe polymer chains resume an amorphous, completely random orientation.

The invention provides for deforming a polymer construct and heatsetting at a temperature that allows the relatively high rate ofnucleation coupled with the relatively low crystal growth rate (See R inFIG. 2 as a general indication of the temperature range at whichdeformation and heat setting of the polymer construct according to theinvention takes place). Thus, the invention provides a method to controlthe relationship between degree of nucleation and associated crystalgrowth rate. At such temperature range, segmental motion of the polymerincreases and polymer chains are able to orient and crystallize, andform a larger number of smaller-sized crystalline domains throughout thepolymer construct.

During radially expansion and heat setting the polymer construct,crystallization is induced. Depending on the application of the polymerconstruct, a selected period of time to heat set the radially expandedpolymer tube may be adjusted to achieve a given crystallinity. Further,the temperature at which the polymer construct is deformed and heat setmay be varied within the temperature range according to the desiredapplication of the polymer construct. Therefore, deformation and heatsetting are performed in a temperature range that allows a larger numberof small crystals, i.e., high nucleation rate and a low crystal growthrate, to achieve a desired circumferential strength and toughness forthe resultant stent.

The invention provides for deforming a polymer construct at atemperature range of from about Tg to about 0.6(Tm−Tg)+Tg, andmaintaining the deformed polymer construct in the deformed state at sucha temperature range for a sufficient period of time to heat set thepolymer construct. By “sufficient period of time” it is meant the timenecessary to form an implantable medical device that can support a lumenin an expanded state for a treatment period. The resultant polymerconstruct yielded from this process has a crystallinity that issubstantially formed of smaller crystalline domains. In addition, thenumber of crystalline domains in a polymer construct according to theinvention increases.

Further, the polymer construct can be a blend of two or more polymers.The polymer construct may be, but is not limited to, a tube, sheet,fiber, or any other shape for use in fabricating an implantable medicaldevice.

At a temperature of about Tg to about 0.6(Tm−Tg)+Tg, the polymerconstruct may be deformed and heat set at a temperature that allows arelatively greater nucleation rate A coupled with a relatively lowercrystal growth rate B. Thus, the invention provides a method to controlthe relationship between degree of nucleation and associated crystalgrowth. At such temperature range, segmental motion of the polymerincreases and polymer chains are able to orient and crystallize, forminga larger number of smaller-sized crystalline domains throughout thepolymer construct. Consequently, physical aging, creep, and stressrelaxation are reduced or curtailed if not eliminated. The relativelylarge number of small crystalline domains inhibits movement of chains inthe amorphous domain.

During deformation of the polymer construct, crystallization is induced.Deforming and heat setting is performed at temperatures near Tm, wherenucleation rate is slow and crystal growth rate is fast, result information of fewer crystals of a larger size. Such a structure tends toexhibit brittleness.

In contrast, the method allows formation of smaller crystals bydeforming and heat setting a polymer construct in a temperature rangewhere nucleation rate A is relatively high, as depicted in FIG. 2, andcrystal growth rate B is relatively low. Deforming and heat setting inthis temperature range forms more crystalline domains of smaller size.Such a structure tends to exhibit less brittleness, and more toughness.

In certain embodiments, a method of fabricating an implantable medicaldevice includes deforming a polymer construct in a temperature range atwhich the crystal nucleation rate A is greater than the crystal growthrate B. In one embodiment, the crystal nucleation rate is greater thanthe crystal growth rate. In another embodiment, the crystal nucleationrate is substantially greater than the crystal growth rate. In oneembodiment, the polymer construct is deformed while maintaining thepolymer construct at a temperature range of from about Tg to about0.6(Tm−Tg)+Tg. In another embodiment, the polymer construct is deformedwhile maintaining the polymer construct at a temperature where the ratioof the crystal nucleation rate to crystal growth rate is 2, 5, 10, 50,100, or greater than 100. In some embodiments, the deformed polymerconstruct is maintained in the deformed state at the selectedtemperature for a sufficient period of time to heat set the polymerconstruct. For example, in one embodiment, the deformed polymerconstruct is maintained at a temperature range of from about Tg to about0.9(Tm−Tg)+Tg. In another embodiment, the deformed polymer construct ismaintained at a temperature where the ratio of the crystal nucleationrate to crystal growth rate is 2, 5, 10, 50, 100, or greater than 100.

The polymer construct of the implantable medical device according to theinvention may have any desired crystallinity according to theapplication of the device. At any given crystallinity, the polymerconstruct has a greater number of smaller sized crystalline domains areformed according to the invention.

FIG. 3A depicts a portion of a strut shown in FIG. 1 of a polymer tubedeformed and heat set at temperatures closer to Tm than Tg according tothe invention.

FIG. 3B depicts a blown up view of a portion 300 of a strut 10 (shown inFIG. 1) that has been deformed and heat set at a temperature closer toTm than Tg, which has fewer larger-sized crystals or crystalline domains310 dispersed in amorphous domain 320, as compared to FIG. 4B, whereportion 400 is deformed and heat set at a temperature closer to Tg thanTm. Portion 300 includes crystals 310 of a larger size becausedeformation and heat setting of polymer construct occurs at atemperature closer to Tm than Tg, where the crystal growth rate issubstantially larger than the nucleation rate. Crystalline domains 310are of a larger size, and thus, fewer crystalline domains are formed. InFIG. 3B, crystalline domains 310 are larger compared to those in portion400. Thus, the regions of the amorphous domain between crystallinedomains in portion 300 are larger.

FIG 4A depicts a portion of a strut shown in FIG. 1 of a polymer tubedeformed and heat set at temperatures closer to Tg than Tm according tothe invention.

FIG. 3B depicts a blown up view of a portion 400 of a strut 110 that hasbeen deformed and heat set at t temperature closer to Tg according tothe invention. Portion 400 has a greater number of smaller crystallinedomains 410, compared with portion 300 that is deformed and heat set ata temperature closer to Tm than Tg. Crystalline domains 410 andamorphous domains 420 of portion 400 are smaller in size. By deformingand heat setting the polymer construct according to the invention for aselected period of time, the size of the crystals can range from lessthan about 10, less than about 6, less than about 2, or less than about1 micron.

Orientation of polymer chains is induced by deformation along the axisof deformation. Small crystalline domains 410 serve as net points toconstrain polymer chains in the amorphous domain 420 of portion 400. Themotion of the polymer chains is restricted through the high number ofsmall crystalline domains 410. A greater number of smaller-sizedcrystals in portion 400 can drastically restrict polymer chain motion inthe amorphous domain 420 while retaining desirable bulk properties ofthe polymer, such as toughness.

Because amorphous domains 420 of portion 400 are constrained by thegreater number of crystalline domains 410, the capacity for movement inthe amorphous domains 420 of portion 400 is decreased. The shorter thedistance between crystalline domains 410, the more the crystallinedomains are able to constrain movement of regions of the amorphousdomain 420. Crystalline domains 410 are able to exert more influence onthe regions of the amorphous domain 420 surrounding crystalline domains410 because the crystalline domains constrain movement of regions of theamorphous domain 420. The crystalline domains 410 or nucleation sitesserve as “tie molecules” within amorphous domain 420 that tie thecrystalline domains together, thereby locking crystalline domains intoplace so that movement of chains in the amorphous domains of the polymerconstruct is reduced. For this reason, portion 400 of the invention hassmaller regions of amorphous domain 420 that resist physical aging,molecule creeping, and stress relaxation of portion 400.

In addition to controlling the physical aging of the polymer construct,the dimensional stability and brittleness of a polymer constructaccording to the invention is controlled. Because crystalline domains ornet points serve as “tie molecules” within the amorphous domain that tiethe crystalline domains together, thereby forming a tighter network ofcrystals that lock the crystalline domains into place, the mechanicalproperties of the polymer construct are improved. For example, bothfracture toughness and shape stability of the polymer construct isincreased. Because a tighter network of crystals is formed, thepossibility of fracturing is substantially reduced because cracks areforced to propagate through and around the many discrete crystals in thepolymer construct. It is believed that the smaller and greater number ofcrystalline domains dispersed throughout the amorphous domain absorbenergy due to crack propagation.

The polymer construct such as a tube, sheet, or fiber, may be deformedusing methods and devices known to persons of skill in the art. Forexample, a polymer tube may be deformed by radially expanding and/oraxially deforming the polymer construct. As discussed above, radialexpansion of the polymer tube can induce circumferential molecularorientation which can increase circumferential strength and modulus orrigidity in the polymer tube. The polymer tube may be expanded radiallyby application of radial pressure. For example, the polymer tube may beexpanded by blow molding.

As mentioned above, since the polymer tube can be expanded and heat setin a temperature range of high nucleation rate and slow crystal growthrate, orientation of the polymer chains induced by radial expansion isbetter maintained by the large number of small crystalline domains. Thepolymer construct is deformed for a sufficient period of time accordingto the desired application of the polymer construct.

In one embodiment, the invention provides for heat setting bymaintaining the deformed polymer construct in the deformed state at atemperature range of from about Tg to about 0.6(Tm−Tg)+Tg for asufficient period of time to heat set the polymer construct. It may bedesirable to heat set the polymer construct at a temperature higher thanthe deformation temperature to increase the rate of the rearrangement ofpolymer chains to adopt to adopt a higher oriented structure. Forexample, in one embodiment, the polymer construct can be heat set at atemperature of about Tg to about 0.9(Tm−Tg)+Tg to allow polymer chainsto adopt the higher oriented structure. The polymer may be maintained inthe deformed state by maintaining a radial pressure and axial tension inthe tube.

After heat setting, the polymer tube may then be cooled to below its Tgeither before or after decreasing the pressure and/or decreasingtension. Cooling the polymer construct helps insure that the tubemaintains the proper shape, size, and length following its formation.Upon cooling, the deformed polymer construct retains the length andshape imposed by an inner surface of the mold.

In one embodiment, the polymer construct consists essentially ofsemi-crystalline poly(L-lactic acid) or “PLLA”. The glass transitiontemperature of PLLA is between 60° C. to about 100° C. The polymerconstruct of PLLA may be made by deforming the polymer construct betweenabout Tg to about 0.6(Tm−Tg)+Tg, and maintaining the deformed polymerconstruct in the deformed state to heat set the polymer construct at atemperature range from about Tg to about 0.9(Tm−Tg)+Tg. In someembodiments, the polymer construct consisting essentially of PLLA isfrom about 45% to about 55% crystalline after deforming and heat settingthe polymer construct. In certain embodiments, the polymer constructconsists essentially of PLLA and is made by deforming the polymerconstruct at a temperature in which the crystal nucleation rate issubstantially greater than the crystal growth rate.

Various embodiments of the polymer construct described above may be usedto fabricate an implantable medical device, such as a stent. Asindicated above, a stent can be formed from a tube or a sheet rolledinto a tube. A sheet or tube, for example, may by formed by variousmethods known in the art such as extrusion or injection blow molding.

Additionally, as indicated above, a stent fabricated from embodiments ofthe polymer construct described herein can be medicated with a bioactiveagent. As used herein, a bioactive agent refers any substance that is ofmedical or veterinary therapeutic, prophylactic or diagnostic utility.Therapeutic use refers to a bioactive agent that, when administered to apatient, will cure, or at least relieve to some extent one or moresymptoms of a disease or disorder. Prophylactic use refers to abioactive agent that, when administered to a patient either prevents theoccurrence of a disease or disorder or, if administered subsequent to atherapeutic agent, prevents or retards the recurrence of the disease ordisorder. For the purposes of this invention, any such agent may beincluded in the construct that is subjected to the method so long as theconditions of the method will not adversely affect the agent.

This invention has been described in relation to certain examples of itsapplication, such as its applicability to constructs comprisingsemi-crystalline PLLA. The examples are not intended nor should they beconstrued as limiting this invention in any manner. Those skilled in theart will recognize, based on the disclosures herein, other polymer andother constructs to which the invention herein may be applied. All suchpolymers and constructs are within the scope of this invention.

What is claimed is:
 1. A method for fabricating an implantable medicaldevice, the method comprising: radially expanding a polymer tube byapplying stress while maintaining the polymer tube at a temperaturerange of from Tg to 0.6(Tm−Tg)+Tg; maintaining the expanded polymer tubein the expanded state at a temperature range of from Tg to 0.9(Tm−Tg)+Tgfor a sufficient period of time to heat set the polymer tube; andfabricating the implantable medical device which is a stent from theheat set polymer tube, wherein the fabricating comprises forming apattern in the heat set polymer tube by cutting which forms ascaffolding including interconnecting structural elements.
 2. The methodaccording to claim 1, wherein the expansion of the polymer tube inducesorientation of the molecules and increases the strength of the polymertube along a direction of the expansion.
 3. The method according toclaim 1, wherein the polymer tube comprises a semi-crystalline polymer.4. The method according to claim 1, wherein the polymer tube comprisespoly(L-lactic acid), and expanding and maintaining is done at atemperature of from 60° C. to 100° C.
 5. The method according to claim1, wherein the heat set polymer tube comprises a crystalline structurehaving crystals at a size less than 2microns or about 2 microns.
 6. Themethod according to claim 1, wherein the poly(L-lactic acid) of thepolymer tube is from about 45% to about 55% crystalline after the radialexpansion and the maintaining of the polymer tube.
 7. A method forfabricating an implantable medical device comprising: radially expandinga polymer tube at a first temperature greater than a Tg of the polymer;wherein the polymer tube is made of poly(L-lactic acid), wherein theradial expanding increases strength of the polymer tube, inducescrystallization in the polymer tube, and induces circumferentialorientation in the polymer tube; maintaining the radially expandedpolymer tube in the expanded state at a second temperature greater thanthe Tg of the polymer for a time period sufficient for a stent formedfrom the polymer tube to support a lumen in a deployed state, whereinthe first temperature during the radial expanding and the secondtemperature during the maintaining are controlled such that the crystalnucleation rate of the polymer tube is greater than the crystal growthrate which maintains the increased strength and induced orientation andresults in small sized crystallite domains that increase fracturetoughness of the polymer tube; and fabricating the stent from thepolymer tube following the maintaining, wherein the fabricatingcomprises forming a pattern in the polymer tube by cutting which forms ascaffolding including interconnecting structural elements.
 8. The methodaccording to claim 7, further comprising axially deforming the polymertube.
 9. The method according to claim 7, wherein the ratio of thecrystal nucleation rate to crystal growth rate is greater than
 2. 10.The method according to claim 7, wherein the first and/or the secondtemperature is in a range of from Tg 0.6(Tm −Tg)+Tg.
 11. The methodaccording to claim 7, wherein the first and second temperatures arebetween 60° C. and 100° C.
 12. The method according to claim 7, whereinthe polymer tube following the maintaining and the stent comprise acrystalline structure having crystals at a size from about 2 to about 10microns.
 13. The method according to claim 7, wherein the polymer tubeis from about 45% to about 55% crystalline after the maintaining.
 14. Amethod for fabricating an implantable medical device, the methodcomprising: radially expanding a polymer tube by applying stress whilemaintaining the polymer tube at a temperature range of from Tg to0.6(Tm−Tg)+Tg; maintaining the expanded polymer tube in the expandedstate at a temperature range of from Tg to 0.9(Tm−Tg)+Tg for a selectedperiod of time, wherein the polymer construct after the radial expandingand the maintaining has crystalline domains less than 10 microns; andfabricating the implantable medical device which is a stent from theheat set polymer tube.